Biocorrosion

[23] Edeleanu C. In: Shrier LL, Jarman RA, Burstein GT, editors. Corrosion. London, Oxford: Butterworth-Heinemann; 1994

Science. 1970;**10**:435

*Corrosion*

**142**

[24] Bockris JO'M, Subramanyan PK. Contributions to the electrochemical basis of the stability of metals. Corrosion

[25] Taleb A, Stafiej J, Badiali JP.

[26] Janik-Czachor M, Wood GC, Thompson GE. Assessment of the processes leading to pit nucleation. British Corrosion Journal. 1980;**15**:154

[27] Shewmon PC. Diffusion in Solids. New York: McGraw-Hill; 1963. p. 175

[28] Hull D. Introduction to Dislocations. Oxford: Pergamon Press; 1968. p. 21

Numerical simulation of metal corrosion with cluster formation. Transactions on Engineering Series. 2005;**48**:109

**Chapter 9**

**Abstract**

alloys.

**145**

**1. Introduction**

treatment (HTT), hydroxyapatite (HA)

Production of Hydroxyapatite on

the Surface of Ti6Al7Nb Alloy as

*Elinor Nahum, Svetlana Lugovskoy and Alex Lugovskoy*

Ti6Al4V is very commonly used for the production of dental implants. Titanium alloys whose mechanical and corrosion properties are equal or better than those of Ti6Al4V might present interest as plausible future materials, too. Ti6Al7Nb alloy was tested and compared to Ti6Al4V in this work. Samples of both alloys were oxidized in a water solution containing calcium acetate (Ca(CH3COO)2) and calcium glycerophosphate (Ca(PO4CH(CH2OH)2)) by Plasma Electrolytic Oxidation (PEO) for 20 min. After that, the samples were hydrothermally treated (HTT) in water (pH = 7) and in potassium hydroxide (KOH) solution (pH = 11) for 2 hours at 200°C in a pressurized reactor. The content and morphology of hydroxyapatite (HA) layers formed on the surface of both alloys after the PEO and subsequent HTT treatments were studied. The surface morphologies, elemental composition, and phase components were characterized by Scanning Electron Microscopy (SEM), Energy Dispersive Spectroscopy (EDS), and X-Ray Diffraction (XRD), respectively. The surface roughness was measured by Atomic Force Microscope (AFM), and thickness measurements were made by SEM and thickness gauge. Corrosion measurements were performed for the comparison of the corrosion behavior of the two

**Keywords:** Ti6Al4V, Ti6Al7Nb, plasma electrolytic oxidation (PEO), hydrothermal

Titanium alloys are often used for the production of various tools or devices to be implanted into a human body: artificial joints, blood vessel prostheses, dental implants, and so on. Of the most popular titanium alloys in that field are Ti6Al4V (Titanium grade 5) and Ti6Al4V-ELI (Titanium grade 23), which both have relatively low Young moduli, compatible with that of the bone issues, good fatigue strength, and excellent corrosion resistance in physiological environments [1]. A layer containing mainly Titania (TiO2) is formed spontaneously on the surface of Titanium alloys. Not only does this layer protect the alloy against corrosion, but it

Other titanium alloys having suitable properties might present both theoretical

also favors their integration with living tissues, that is, *osseointegration* [2].

and applied interest as the novel materials for medical device production. A Niobium-containing Ti6Al7Nb is one of such alloys. The corrosion behavior of

Compared to Ti6Al4V Alloy

#### **Chapter 9**

## Production of Hydroxyapatite on the Surface of Ti6Al7Nb Alloy as Compared to Ti6Al4V Alloy

*Elinor Nahum, Svetlana Lugovskoy and Alex Lugovskoy*

#### **Abstract**

Ti6Al4V is very commonly used for the production of dental implants. Titanium alloys whose mechanical and corrosion properties are equal or better than those of Ti6Al4V might present interest as plausible future materials, too. Ti6Al7Nb alloy was tested and compared to Ti6Al4V in this work. Samples of both alloys were oxidized in a water solution containing calcium acetate (Ca(CH3COO)2) and calcium glycerophosphate (Ca(PO4CH(CH2OH)2)) by Plasma Electrolytic Oxidation (PEO) for 20 min. After that, the samples were hydrothermally treated (HTT) in water (pH = 7) and in potassium hydroxide (KOH) solution (pH = 11) for 2 hours at 200°C in a pressurized reactor. The content and morphology of hydroxyapatite (HA) layers formed on the surface of both alloys after the PEO and subsequent HTT treatments were studied. The surface morphologies, elemental composition, and phase components were characterized by Scanning Electron Microscopy (SEM), Energy Dispersive Spectroscopy (EDS), and X-Ray Diffraction (XRD), respectively. The surface roughness was measured by Atomic Force Microscope (AFM), and thickness measurements were made by SEM and thickness gauge. Corrosion measurements were performed for the comparison of the corrosion behavior of the two alloys.

**Keywords:** Ti6Al4V, Ti6Al7Nb, plasma electrolytic oxidation (PEO), hydrothermal treatment (HTT), hydroxyapatite (HA)

#### **1. Introduction**

Titanium alloys are often used for the production of various tools or devices to be implanted into a human body: artificial joints, blood vessel prostheses, dental implants, and so on. Of the most popular titanium alloys in that field are Ti6Al4V (Titanium grade 5) and Ti6Al4V-ELI (Titanium grade 23), which both have relatively low Young moduli, compatible with that of the bone issues, good fatigue strength, and excellent corrosion resistance in physiological environments [1]. A layer containing mainly Titania (TiO2) is formed spontaneously on the surface of Titanium alloys. Not only does this layer protect the alloy against corrosion, but it also favors their integration with living tissues, that is, *osseointegration* [2].

Other titanium alloys having suitable properties might present both theoretical and applied interest as the novel materials for medical device production. A Niobium-containing Ti6Al7Nb is one of such alloys. The corrosion behavior of

Ti6Al7Nb in the simulated body fluid (SBF) was studied by Rajendran et al. and was found comparable or better than that of Ti6Al4V-ELI [3].

electrode. All the corrosion tests were performed in Hank's solution [10] and a simulated saliva solution [11], whose chemical compositions are given in **Table 1**. The pH of the electrolytes was 7, and the temperature was maintained at 36.5°C.

*Production of Hydroxyapatite on the Surface of Ti6Al7Nb Alloy as Compared to Ti6Al4V Alloy*

The surface of both titanium alloys after the PEO has the typical for that technique microstructure characterized by microscopic pores scattered randomly across

the surface. Cracks seen on the surface are more pronounced for Ti6Al4V

After the hydrothermal treatment, the surface has changed. If the HT is performed in distilled water at pH = 7, Ti6Al4V surface is characterized by grainy HA crystals on the surface and very small needle-like HA crystals inside the pores (**Figure 2a**). Unlike that, numerous HA platelets are observed both inside and outside the pores on the surface of Ti6Al7Nb (**Figure 2b**). If the HT is made at pH = 11, the surface of Ti6Al4V is covered by ununiform plates of significantly larger HA crystals inside and outside the pores (**Figure 2c**). Ti6Al7Nb surface

**Hank's solution [10] Saliva solution [11]**

*Back scattered electrons SEM images of the sample's surface after PEO, 10,000: (a) Ti6Al4V and*

CaCl2∙2H2O 0.185 MgCl26H2O 0.059 MgSO4 0.09767 KCl 0.625 KCl 0.4 KH2PO4 0.326 KH2PO4 0.06 K2HPO4 0.804 NaHCO3 0.35 CaCl22H2O 0.166 NaCl 8.0 C3H8O3 2.00 Na2HPO4 0.04788 Sodium carboxymethyl cellulose 10.0

**Composition, g/L Reagent Composition, g/L**

**3. Results and discussion**

*DOI: http://dx.doi.org/10.5772/intechopen.92314*

Glucose 1.0

*Chemical composition of Hank's and simulated saliva solutions.*

**Table 1.**

**Figure 1.**

**147**

*(b) Ti6Al7Nb.*

(**Figure 1**).

While the bioinertness (including corrosion stability) of titanium alloys is high enough, their readiness to osseointegration leaves much to be desired [1]. One of the plausible strategies allowing a considerable improvement in the osseointegration of titanium alloys is the production of a layer of Hyfroxyapatite (HA) [4] on their surface. Being a mineral constituent of the bone tissue, Hydroxyapatite is an ideal binder between the metal and the living body.

In this study, the surface modification aiming at the production of HA on the surface of Ti6Al4V and Ti6Al7Nb was made by using Plasma Electrolyte Oxidation (PEO), which is a simple technique for producing hard and rough coating having numerous micro-pores [5, 6]. Using that technique, the insertion into the coating of such elements as calcium and phosphorous may be performed by just adding them to the electrolyte in a suitable form. A PEO layer may also present a diffusional and sorption barrier to the release of metal ions into physiological liquids, thus improving the bioinertness of the core metal [7]. PEO coatings often have good adhesion to the metal even if the implant geometry is complex such as screw-shaped implants [8, 9]. PEO by itself does not cause the growth of HA crystals, rather a specimen needs an additional hydrothermal treatment (HT) [4].

The aim of this study is to compare the efficacy of the production of HA on the surfaces of Ti6Al4V and Ti6Al7Nb by PEO and the subsequent HT.

#### **2. Experimental**

Ti6Al4V samples of 40 mm 20 mm 1 mm size and Ti6Al7Nb samples of 40 mm 20 mm 3 mm size were cut by laser and grounded by 150, 360, 600, and 1000 grid silicon carbide (SiC) papers. The specimens were rinsed in distilled water and acetone in an ultrasonic cleaner for 5 min. The PEO was performed by the 50 Hz sinusoidal AC current in an electrolyte containing 0.25 M calcium acetate and 0.06 M calcium glycerophosphate in distilled water at the current density of 4A/ dm<sup>2</sup> for 20 min. The PEO process occurred in a water-cooled stainless-steel container serving as the counter electrode, equipped with a mechanical stirrer. After the completion of the PEO process, the specimens were washed in distilled water and dried on air. After that, the specimens were hydrothermally treated in distilled water or in a KOH solution at 200°C in a pressurized reactor for 2 hours. The pressure during the treatment was 13–15 bar.

The surface morphology and elemental composition were characterized by scanning microscope electron (SEM) TESCAN MAIA3 TriglavTM equipped with AZteq Oxford energy dispersive spectroscopy (EDS) analyzer. X-Ray diffraction (XRD) Rigaku, SmartLab X-RAY DIFRACTOMETER using Cu-Kα radiation (λ = 1.54 Å) in the range of 15–65° angles with a step 0.02° was used to characterize the phase components of the substrates and coating. The thickness of the coatings was measured by ElektroPhysik MiniTest 730 thickness gauge based on eddy current principle by an average of 10 measurements. Focused ion beam (FIB) technique FEI Helios NanoLab™ 600 DualBeam was used for the production of crosssectional area on a specimen to be further characterized by SEM-EDS. Surface roughness of the samples was evaluated with atomic force microscope (AFM) Bruker's Dimension FastScan with ScanAsystTM using the contact mode.

The corrosion resistance was determined on an IVIUMnSTAT potentiostat by electrochemical polarization methods, namely Linear Polarization Resistance (LPR) and Tafel Slope Extrapolation (TSE) using a three-electrode cell, where an Ag|AgCl electrode served as the reference electrode, and a platinum wire was the counter

*Production of Hydroxyapatite on the Surface of Ti6Al7Nb Alloy as Compared to Ti6Al4V Alloy DOI: http://dx.doi.org/10.5772/intechopen.92314*

electrode. All the corrosion tests were performed in Hank's solution [10] and a simulated saliva solution [11], whose chemical compositions are given in **Table 1**. The pH of the electrolytes was 7, and the temperature was maintained at 36.5°C.

#### **3. Results and discussion**

Ti6Al7Nb in the simulated body fluid (SBF) was studied by Rajendran et al. and was

While the bioinertness (including corrosion stability) of titanium alloys is high enough, their readiness to osseointegration leaves much to be desired [1]. One of the plausible strategies allowing a considerable improvement in the osseointegration of titanium alloys is the production of a layer of Hyfroxyapatite (HA) [4] on their surface. Being a mineral constituent of the bone tissue, Hydroxyapatite is an ideal

In this study, the surface modification aiming at the production of HA on the surface of Ti6Al4V and Ti6Al7Nb was made by using Plasma Electrolyte Oxidation (PEO), which is a simple technique for producing hard and rough coating having numerous micro-pores [5, 6]. Using that technique, the insertion into the coating of such elements as calcium and phosphorous may be performed by just adding them to the electrolyte in a suitable form. A PEO layer may also present a diffusional and sorption barrier to the release of metal ions into physiological liquids, thus improving the bioinertness of the core metal [7]. PEO coatings often have good adhesion to the metal even if the implant geometry is complex such as screw-shaped implants [8, 9]. PEO by itself does not cause the growth of HA crystals, rather a specimen

The aim of this study is to compare the efficacy of the production of HA on the

Ti6Al4V samples of 40 mm 20 mm 1 mm size and Ti6Al7Nb samples of 40 mm 20 mm 3 mm size were cut by laser and grounded by 150, 360, 600, and 1000 grid silicon carbide (SiC) papers. The specimens were rinsed in distilled water and acetone in an ultrasonic cleaner for 5 min. The PEO was performed by the 50 Hz sinusoidal AC current in an electrolyte containing 0.25 M calcium acetate and 0.06 M calcium glycerophosphate in distilled water at the current density of 4A/ dm<sup>2</sup> for 20 min. The PEO process occurred in a water-cooled stainless-steel container serving as the counter electrode, equipped with a mechanical stirrer. After the completion of the PEO process, the specimens were washed in distilled water and dried on air. After that, the specimens were hydrothermally treated in distilled water or in a KOH solution at 200°C in a pressurized reactor for 2 hours. The

The surface morphology and elemental composition were characterized by scanning microscope electron (SEM) TESCAN MAIA3 TriglavTM equipped with AZteq Oxford energy dispersive spectroscopy (EDS) analyzer. X-Ray diffraction (XRD) Rigaku, SmartLab X-RAY DIFRACTOMETER using Cu-Kα radiation (λ = 1.54 Å) in the range of 15–65° angles with a step 0.02° was used to characterize the phase components of the substrates and coating. The thickness of the coatings was measured by ElektroPhysik MiniTest 730 thickness gauge based on eddy current principle by an average of 10 measurements. Focused ion beam (FIB) technique FEI Helios NanoLab™ 600 DualBeam was used for the production of crosssectional area on a specimen to be further characterized by SEM-EDS. Surface roughness of the samples was evaluated with atomic force microscope (AFM) Bruker's Dimension FastScan with ScanAsystTM using the contact mode.

The corrosion resistance was determined on an IVIUMnSTAT potentiostat by electrochemical polarization methods, namely Linear Polarization Resistance (LPR) and Tafel Slope Extrapolation (TSE) using a three-electrode cell, where an Ag|AgCl electrode served as the reference electrode, and a platinum wire was the counter

found comparable or better than that of Ti6Al4V-ELI [3].

binder between the metal and the living body.

needs an additional hydrothermal treatment (HT) [4].

pressure during the treatment was 13–15 bar.

**2. Experimental**

*Corrosion*

**146**

surfaces of Ti6Al4V and Ti6Al7Nb by PEO and the subsequent HT.

The surface of both titanium alloys after the PEO has the typical for that technique microstructure characterized by microscopic pores scattered randomly across the surface. Cracks seen on the surface are more pronounced for Ti6Al4V (**Figure 1**).

After the hydrothermal treatment, the surface has changed. If the HT is performed in distilled water at pH = 7, Ti6Al4V surface is characterized by grainy HA crystals on the surface and very small needle-like HA crystals inside the pores (**Figure 2a**). Unlike that, numerous HA platelets are observed both inside and outside the pores on the surface of Ti6Al7Nb (**Figure 2b**). If the HT is made at pH = 11, the surface of Ti6Al4V is covered by ununiform plates of significantly larger HA crystals inside and outside the pores (**Figure 2c**). Ti6Al7Nb surface


#### **Table 1.**

*Chemical composition of Hank's and simulated saliva solutions.*

**Figure 1.**

*Back scattered electrons SEM images of the sample's surface after PEO, 10,000: (a) Ti6Al4V and (b) Ti6Al7Nb.*

hydroxyapatite is 1.67 that was not observed for any specimen, which means that the surfaces always contain mixtures of various calcium phosphates rather than the

*Elemental composition (at%, EDS) of the surfaces after PEO and hydrothermal treatments.*

*A typical FIB-ablated cross-sectional structure of the surface: (a) titanium alloy substrate, (b) PEO porous*

*Production of Hydroxyapatite on the Surface of Ti6Al7Nb Alloy as Compared to Ti6Al4V Alloy*

Ti6Al4V, PEO 15.1 0.1 1.6 0.0 0.6 0.0 63.3 0.2 8.6 0.0 5.7 0.0 1.51 Ti6Al7Nb, PEO 10.5 0.5 1.2 0.0 0.4 0.0 58.1 0.6 7.5 0.1 4.6 0.0 1.63 Ti6Al4V, HT pH = 7 17.7 1.9 1.9 0.0 3.0 0.2 64.8 0.6 7.2 0.4 4.4 0.2 1.64 Ti6Al7Nb, HT pH = 7 17.9 0.3 2.2 0.1 0.7 0.1 69.8 0.0 5.0 0.3 4.6 0.1 1.09 Ti6Al4V, HT pH = 11 15.5 4.6 1.6 0.7 0.4 0.1 59.9 7.7 11.6 8.5 5.8 3.4 2.00 Ti6Al7Nb, HT pH = 11 14.5 3.4 2.1 0.6 0.6 0.0 65.3 5.3 6.1 3.7 5.1 1.9 1.20

**Ti Al Nb or V O Ca P Ca/P**

The thickness of the coating was determined by two different methods (**Table 4**), namely by using a thickness gauge and measuring the FIB-ablated cross sections in SEM images. The measurements reveal higher coating thicknesses for Ti6Al4V than Ti6Al7Nb. Ti6Al4V coating shows also larger and more uneven

3D AFM images of Ti6Al4V and Ti6Al7Nb after PEO and hydrothermal treatments are shown in **Figure 5**. The area scanned was 5 μm 5 μm, and three sites were scanned for each specimen. The values of average roughness (*R*a) for all the

In order to determine the phase composition of the surfaces, XRD spectra were measured (**Figure 4** and **Table 3**). It can be seen from **Figure 4** and **Table 3** that for both alloys no detectable amount of HA is present after PEO. Rather, the surfaces are covered by the mixture of rutile and anatase. Additionally, the surface of Ti6Al4V contains a small amount of tricalcium phosphate Ca2(PO4)2, which is not the case for Ti6Al7Nb. After the HT treatment, an HA phase is detected on the

pure hydroxyapatite.

**Figure 3.**

**Table 2.**

**149**

*layer, and (c) hydroxyapatite layer.*

*DOI: http://dx.doi.org/10.5772/intechopen.92314*

surface of both alloys.

thickness for Ti6Al4V than for Ti6Al7Nb.

#### **Figure 2.**

*Surface morphologies after hydrothermal treatment, BSE SEM 10,000; 30,000 in the inserts: (a) Ti6Al4V pH = 7; (b) Ti6Al7Nb pH = 7; (c) Ti6Al4V pH = 11; and (d) Ti6Al7Nb pH = 11.*

contains large plates of HA inside the pores and a mixture of grainy and needle-like crystals outside the pores (**Figure 2d**).

After the completion of the PEO + HT treatment, the surface layers were partially ablated by FIB, so that the 'cross-sectional'structure could be seen (**Figure 3**). As is seen in **Figure 3**, there is an approximately 1 μm porous PEO oxide layer on the surface of the alloys. The oxide layer has partially amorphous and partially fine crystalline structure (region 'b' in the image); above that, an approximately 1 μm hydroxyapatite layer (region 'c' in the image) consisting of larger crystallites is present. The thicknesses of the oxide and hydroxyapatite layers may vary from one specimen to another, while their structure remains the same.

The elemental compositions of the surfaces obtained by EDS are given in **Table 2**. The presented chemical composition is the average of three-point mode analysis, and standard deviations are displayed. The stoichiometric Ca/P ratio for *Production of Hydroxyapatite on the Surface of Ti6Al7Nb Alloy as Compared to Ti6Al4V Alloy DOI: http://dx.doi.org/10.5772/intechopen.92314*

#### **Figure 3.**

*A typical FIB-ablated cross-sectional structure of the surface: (a) titanium alloy substrate, (b) PEO porous layer, and (c) hydroxyapatite layer.*


#### **Table 2.**

contains large plates of HA inside the pores and a mixture of grainy and needle-like

*Surface morphologies after hydrothermal treatment, BSE SEM 10,000; 30,000 in the inserts: (a) Ti6Al4V*

After the completion of the PEO + HT treatment, the surface layers were partially ablated by FIB, so that the 'cross-sectional'structure could be seen (**Figure 3**). As is seen in **Figure 3**, there is an approximately 1 μm porous PEO oxide layer on the surface of the alloys. The oxide layer has partially amorphous and partially fine crystalline structure (region 'b' in the image); above that, an approximately 1 μm hydroxyapatite layer (region 'c' in the image) consisting of larger crystallites is present. The thicknesses of the oxide and hydroxyapatite layers may vary from one

The elemental compositions of the surfaces obtained by EDS are given in **Table 2**. The presented chemical composition is the average of three-point mode analysis, and standard deviations are displayed. The stoichiometric Ca/P ratio for

crystals outside the pores (**Figure 2d**).

**Figure 2.**

*Corrosion*

**148**

specimen to another, while their structure remains the same.

*pH = 7; (b) Ti6Al7Nb pH = 7; (c) Ti6Al4V pH = 11; and (d) Ti6Al7Nb pH = 11.*

*Elemental composition (at%, EDS) of the surfaces after PEO and hydrothermal treatments.*

hydroxyapatite is 1.67 that was not observed for any specimen, which means that the surfaces always contain mixtures of various calcium phosphates rather than the pure hydroxyapatite.

In order to determine the phase composition of the surfaces, XRD spectra were measured (**Figure 4** and **Table 3**). It can be seen from **Figure 4** and **Table 3** that for both alloys no detectable amount of HA is present after PEO. Rather, the surfaces are covered by the mixture of rutile and anatase. Additionally, the surface of Ti6Al4V contains a small amount of tricalcium phosphate Ca2(PO4)2, which is not the case for Ti6Al7Nb. After the HT treatment, an HA phase is detected on the surface of both alloys.

The thickness of the coating was determined by two different methods (**Table 4**), namely by using a thickness gauge and measuring the FIB-ablated cross sections in SEM images. The measurements reveal higher coating thicknesses for Ti6Al4V than Ti6Al7Nb. Ti6Al4V coating shows also larger and more uneven thickness for Ti6Al4V than for Ti6Al7Nb.

3D AFM images of Ti6Al4V and Ti6Al7Nb after PEO and hydrothermal treatments are shown in **Figure 5**. The area scanned was 5 μm 5 μm, and three sites were scanned for each specimen. The values of average roughness (*R*a) for all the

osseointegration because it is compatible with the sizes of small cells and large

For the determination of corrosion parameters of the alloys, polarization curves of the specimens in Hank's solution and in artificial saliva were measured in the range of 250 mV with respect to the OCP at the scan rate of 1 mV/s. Additionally, linear polarization measurements (LPRs) were performed in the narrower range of 10 mV at the scan rate of 0.5 mV/s. The measured values of corrosion current densities and corrosion potentials are given in **Table 6**. For some cases (these are marked gray in **Table 6**), it was not possible to measure the corrosion parameters

8.5 1.0 8.8 1.1

7.0 0.9 9.5 1.2

3.2 0.1 2.4 0.2

3.0 0.6 2.6 0.4

As is seen from **Table 6**, all the corrosion potentials that could be measured are significantly shifted to more noble values after the hydrothermal treatment, so that the alloy is effectively passivated. No essential difference was observed for the corrosion potential of the two alloys (at least, when those were measurable).

The values of corrosion current densities are scattered in a quite random manner

and therefore are less informative. We assume that due to the relatively poor electrical conductivity of both liquids (the WC-CE resistance measured in the cell was 10–20 kΩ for Hank's solution and for the artificial saliva, which is at least by the factor of 1000 higher than for such strong electrolytes as KCl), the precision of

The morphologies, elemental and phase's composition, coating thickness, roughness, and corrosion behavior of Ti6Al4V and Ti6Al7Nb alloys after Plasma Electrolytic Oxidation and the subsequent hydrothermal treatment at various pHs were studied and compared. Hydroxyapatite-containing surfaces can be attained by

Thicker, finer, and more uniform oxide layers are formed on the surface of

The most developed surface with plate-shaped HA crystals was obtained for

The corrosion potentials are significantly shifted to more noble values after the hydrothermal treatment, so that the alloy is effectively passivated. No essential difference was observed between the corrosion potential of the two alloys. It was

biomolecules [12, 13].

**Table 4.**

**4. Conclusions**

**151**

because the systems were too passive.

**Ti alloy Treatment Average coating thickness by**

*DOI: http://dx.doi.org/10.5772/intechopen.92314*

HTT pH = 7

HTT pH = 11

HTT pH = 7

HTT pH = 11

*Thickness of the coatings on Ti alloys.*

**SEM, μm**

Ti6Al4V PEO 6.4 1.2 9.4 1.0

*Production of Hydroxyapatite on the Surface of Ti6Al7Nb Alloy as Compared to Ti6Al4V Alloy*

Ti6Al7Nb PEO 2.8 0.4 2.8 0.3

**Average coating thickness by thickness gauge, μm**

the polarization methods was not sufficient.

Ti6Al7Nb after HTT in distilled water.

the two-stage procedure, PEO and HTT, for both alloys.

Ti6Al4V than on Ti6Al7Nb for the same treatment parameters.

**Figure 4.** *XRD spectra acquired from the coatings: (a) Ti-6Al-4 V and (b) Ti-6Al-7Nb.*


#### **Table 3.**

*Phase composition of coating after PEO and hydrothermal treatments.*

specimens lie in the range of 50–250 nm (**Table 5**), and for Ti6Al7Nb, they are higher for all the treatments. As is seen from both AFM (**Figure 5**) and SEM (**Figure 2**) images, a more developed surface with plate-shaped HA crystals inside the pores and grainy crystals on the surface is formed on Ti6Al7Nb than on Ti6Al4V. The roughness range of 10 nm to 10 μm is favorable for the

*Production of Hydroxyapatite on the Surface of Ti6Al7Nb Alloy as Compared to Ti6Al4V Alloy DOI: http://dx.doi.org/10.5772/intechopen.92314*


#### **Table 4.**

*Thickness of the coatings on Ti alloys.*

osseointegration because it is compatible with the sizes of small cells and large biomolecules [12, 13].

For the determination of corrosion parameters of the alloys, polarization curves of the specimens in Hank's solution and in artificial saliva were measured in the range of 250 mV with respect to the OCP at the scan rate of 1 mV/s. Additionally, linear polarization measurements (LPRs) were performed in the narrower range of 10 mV at the scan rate of 0.5 mV/s. The measured values of corrosion current densities and corrosion potentials are given in **Table 6**. For some cases (these are marked gray in **Table 6**), it was not possible to measure the corrosion parameters because the systems were too passive.

As is seen from **Table 6**, all the corrosion potentials that could be measured are significantly shifted to more noble values after the hydrothermal treatment, so that the alloy is effectively passivated. No essential difference was observed for the corrosion potential of the two alloys (at least, when those were measurable).

The values of corrosion current densities are scattered in a quite random manner and therefore are less informative. We assume that due to the relatively poor electrical conductivity of both liquids (the WC-CE resistance measured in the cell was 10–20 kΩ for Hank's solution and for the artificial saliva, which is at least by the factor of 1000 higher than for such strong electrolytes as KCl), the precision of the polarization methods was not sufficient.

#### **4. Conclusions**

The morphologies, elemental and phase's composition, coating thickness, roughness, and corrosion behavior of Ti6Al4V and Ti6Al7Nb alloys after Plasma Electrolytic Oxidation and the subsequent hydrothermal treatment at various pHs were studied and compared. Hydroxyapatite-containing surfaces can be attained by the two-stage procedure, PEO and HTT, for both alloys.

Thicker, finer, and more uniform oxide layers are formed on the surface of Ti6Al4V than on Ti6Al7Nb for the same treatment parameters.

The most developed surface with plate-shaped HA crystals was obtained for Ti6Al7Nb after HTT in distilled water.

The corrosion potentials are significantly shifted to more noble values after the hydrothermal treatment, so that the alloy is effectively passivated. No essential difference was observed between the corrosion potential of the two alloys. It was

specimens lie in the range of 50–250 nm (**Table 5**), and for Ti6Al7Nb, they are higher for all the treatments. As is seen from both AFM (**Figure 5**) and SEM (**Figure 2**) images, a more developed surface with plate-shaped HA crystals inside the pores and grainy crystals on the surface is formed on Ti6Al7Nb than on Ti6Al4V. The roughness range of 10 nm to 10 μm is favorable for the

**Treatment Ti6Al7Nb Ti6Al4V**

Amorphous phase

HA

HA

TiO2-rutile, anatase Ca3(PO4)2 Amorphous phase

TiO2-rutile, anatase HA

TiO2-rutile, anatase HA

*XRD spectra acquired from the coatings: (a) Ti-6Al-4 V and (b) Ti-6Al-7Nb.*

PEO TiO2-rutile, anatase

HTT pH = 7 TiO2-anatase

HTT pH = 11 TiO2-anatase

*Phase composition of coating after PEO and hydrothermal treatments.*

**Figure 4.**

*Corrosion*

**Table 3.**

**150**

found that the polarization corrosion measurement is not precise enough for both alloys in Hank's solution and in the artificial saliva because of the poor conductivity

**Ecorr, mV vs***.* **Ag|AgCl jcorr, A/cm<sup>2</sup>**

Saliva Passive

Saliva <sup>303</sup> 2.91<sup>10</sup><sup>8</sup> 2.63<sup>10</sup><sup>8</sup>

Saliva <sup>156</sup> 2.06<sup>10</sup><sup>7</sup> 1.02<sup>10</sup><sup>7</sup>

Saliva <sup>127</sup> 2.32<sup>10</sup><sup>7</sup> 1.30<sup>10</sup><sup>7</sup>

Ti6Al4V, PEO Hank'<sup>s</sup> <sup>293</sup> 3.76<sup>10</sup><sup>7</sup> 2.31<sup>10</sup><sup>7</sup>

*Production of Hydroxyapatite on the Surface of Ti6Al7Nb Alloy as Compared to Ti6Al4V Alloy*

Ti6Al7Nb, PEO Hank'<sup>s</sup> <sup>403</sup> 2.66<sup>10</sup><sup>6</sup> 1.32<sup>10</sup><sup>6</sup>

Ti6Al4V, HT pH = 7 Hank'<sup>s</sup> <sup>151</sup> 5.16<sup>10</sup><sup>7</sup> 2.83<sup>10</sup><sup>7</sup>

Saliva Passive Ti6Al4V, HT pH = 11 Hank'<sup>s</sup> <sup>112</sup> 1.38<sup>10</sup><sup>6</sup> 1.23<sup>10</sup><sup>6</sup>

Saliva Passive

*Corrosion current density and corrosion potentials for Ti6Al4V and Ti6Al7Nb in Hank's and saliva solutions.*

Ti6Al7Nb, HT pH = 7 Hank's Passive

*DOI: http://dx.doi.org/10.5772/intechopen.92314*

Ti6Al7Nb, HT pH = 11 Hank's Passive

**, LPR jcorr, A/cm2**

**, Tafel**

of both liquids.

**Table 6.**

**Author details**

**153**

Elinor Nahum\*, Svetlana Lugovskoy and Alex Lugovskoy

\*Address all correspondence to: elinorna@ariel.ac.il

provided the original work is properly cited.

Department of Chemical Engineering, Ariel University, Ariel, Israel

© 2020 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium,

**Figure 5.**

*3D AFM images: (a) Ti6Al4V after PEO; (b) Ti6Al4V after HTT pH = 7; (c) Ti6Al4V after HTT pH = 11; (d) Ti6Al7Nb after PEO; (e) Ti6Al7Nb after HTT pH = 7; (f) Ti6Al7Nb after HTT pH = 11.*


#### **Table 5.**

R*<sup>a</sup> values (nm) for Ti6Al4V and Ti6Al7Nb surfaces after PEO and HT treatments.*

*Production of Hydroxyapatite on the Surface of Ti6Al7Nb Alloy as Compared to Ti6Al4V Alloy DOI: http://dx.doi.org/10.5772/intechopen.92314*


#### **Table 6.**

*Corrosion current density and corrosion potentials for Ti6Al4V and Ti6Al7Nb in Hank's and saliva solutions.*

found that the polarization corrosion measurement is not precise enough for both alloys in Hank's solution and in the artificial saliva because of the poor conductivity of both liquids.

#### **Author details**

Elinor Nahum\*, Svetlana Lugovskoy and Alex Lugovskoy Department of Chemical Engineering, Ariel University, Ariel, Israel

\*Address all correspondence to: elinorna@ariel.ac.il

© 2020 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

**PEO HT, pH = 7 HT, pH = 11**

Ti6Al4V 83.1 115.7 55.4 Ti6Al7Nb 232.7 175.3 234.7

*(d) Ti6Al7Nb after PEO; (e) Ti6Al7Nb after HTT pH = 7; (f) Ti6Al7Nb after HTT pH = 11.*

*3D AFM images: (a) Ti6Al4V after PEO; (b) Ti6Al4V after HTT pH = 7; (c) Ti6Al4V after HTT pH = 11;*

R*<sup>a</sup> values (nm) for Ti6Al4V and Ti6Al7Nb surfaces after PEO and HT treatments.*

**Table 5.**

**152**

**Figure 5.**

#### **References**

[1] Manivasagam G, Dhinasekaran D, Rajamanickam A. Biomedical implants: Corrosion and its prevention - A review. Recent Patents on Corrosion Science. 2010;**2**:40-54. DOI: 10.2174/ 1877610801002010040

[2] Jäger M, Jennissen HP, Dittrich F, Fischer A, Köhling HL. Antimicrobial and osseointegration properties of nanostructured titanium orthopaedic implans. Materials. 2017;**10**:1-28. DOI: 10.3390/ma10111302

[3] Tamilselvi S, Raman V, Rajendran N. Corrosion behaviour of Ti-6Al-7Nb and Ti-6Al-4V ELI alloys in the simulated body fluid solution by electrochemical impedance spectroscopy. Electrochimica Acta. 2006;**52**:839-846. DOI: 10.1016/j.electacta.2006.06.018

[4] Lugovskoy A, Lugovskoy S. Production of hydroxyapatite layers on the plasma electrolytically oxidized surface of titanium alloys. Materials Science and Engineering: C. 2014;**43**: 527-532. DOI: 10.1016/j. msec.2014.07.030

[5] Lugovskoy A, Zinigrad M. Plasma electrolytic oxidation of valve metals. In: Mastai Y, editor. Advances in Materials Science and Engineering. London: InTech; 2013. pp. 85-102. DOI: 10.5772/ 54827

[6] Ibrahim MZ, Sarhan AAD, Yusuf F, Hamdi M. Biomedical materials and techniques to improve the tribological, mechanical and biomedical properties of orthopedic implants – A review article. Journal of Alloys and Compounds. 2017; **714**:636-667. DOI: 10.1016/j. jallcom.2017.04.231

[7] Mohedano M, Matykina E, Arrabal R, Pardo A, Merino MC. Metal release from ceramic coatings for dental implants. Dental Materials. 2014;**30**:e28-e40. DOI: 10.1016/j.dental.2013.12.011

[8] Zhu X, Chen J, Scheideler L, Reichl R, Geis-Gerstorfer J. Effects of topography and composition of titanium surface oxides on osteoblast responses. Biomaterials. 2004;**25**:4087-4103. DOI: 10.1016/j.biomaterials.2003.11.011

[9] Esen Z, Öcal EB. Surface characteristics and in-vitro behavior of chemically treated bulk Ti6Al7Nb alloys. Surface and Coatings Technology. 2017;**309**:829-839. DOI: 10.1016/j.surfcoat.2016.10.078

[10] Balanced Hank's Solution. Available from: https://www.sigmaaldrich.com/c ontent/dam/sigma-aldrich/docs/Sigma/ Formulation/d5796for.pdf

[11] Simulated Saliva Solution. Available from: https://www.pickeringlabs.c om/wp-content/uploads/sds/SDS 1700 0305 Artificial Saliva for Medical and Dental Research.pdf

[12] Dilea M, Mazare A, Ionita D, Demetrescu I. Comparison between corrosion behaviour of implant alloys Ti6Al7Nb and Ti6Al4Zr in artificial saliva. Materials and Corrosion. 2013; **64**:493-499. DOI: 10.1002/ maco.201206526

[13] Elias CN, Meirelles L. Improving osseointegration of dental implants. Expert Review of Medical Devices. 2010;**7**:241-256. DOI: 10.1586/erd.09.74

**155**

**Chapter 10**

**Abstract**

counterbody.

**1. Introduction**

wear volume, coefficient of friction

[22], where *R* is the radius of the ball.

*<sup>V</sup>* <sup>≅</sup> *<sup>π</sup> <sup>d</sup>* \_

mechanical structure with relative movement.

Biotribology of Mechanically and

*Jorge Humberto Luna-Domínguez and Ronaldo Câmara Cozza*

The purpose of present work is to study the biotribological behavior of a mechanically and laser marked biomaterial. Sliding wear tests were conducted on ASTM F139 austenitic stainless-steel specimen, with polypropylene and AISI 316 L austenitic stainless-steel balls, as counterbodies. During wear experiments, a liquid chemical composition was continuously fed between the specimen and the ball. The coefficient of friction acting on the tribological system "specimen – liquid chemical composition – ball" and the wear volume of the wear craters were calculated, and results were analyzed. The results have shown that the biotribological behavior of ASTM F139 austenitic stainless steel was influenced by mechanical or laser marking process, and its wear resistance was dependent on the kind of

**Keywords:** biomaterial, austenitic stainless steel, laser treatment, wear resistance,

The "ball-cratering wear test" has gained large acceptance at universities and research centers as it is widely used in studies focusing on the wear behavior of different materials [1–20]. **Figure 1** presents a schematic diagram of the principle of wear test, where a rotating ball is forced against the tested specimen and liquid solution supplied between the specimen and the ball during the experiments. The aim of the "ball-cratering wear test" is to generate "wear craters" on the surface of the specimen. **Figure 2** presents an image of such crater, together with an indication of the crater diameter (*d*) (**Figure 2a**) and the wear volume (*V*) (**Figure 2b** [21]). The wear volume is determined as a function of "*d*," using Eq. (1)

4

In other line of research, the concept of "biotribology" has gained important spotlight in the area, including research works addressing the biotribological behavior of materials [23–29] used in the manufacturing of human body elements. Consequently, different laboratory techniques have been employed to reproduce conditions where there are friction and consequent wear of parts of the human

<sup>64</sup>*<sup>R</sup>* for *<sup>d</sup>* << *<sup>R</sup>* (1)

Laser Marked Biomaterial

*Marcelo de Matos Macedo, Vikas Verma,* 

#### **Chapter 10**

**References**

*Corrosion*

[1] Manivasagam G, Dhinasekaran D, Rajamanickam A. Biomedical implants: Corrosion and its prevention - A review. Recent Patents on Corrosion Science.

[8] Zhu X, Chen J, Scheideler L, Reichl R, Geis-Gerstorfer J. Effects of topography and composition of titanium surface oxides on osteoblast responses. Biomaterials. 2004;**25**:4087-4103. DOI: 10.1016/j.biomaterials.2003.11.011

[9] Esen Z, Öcal EB. Surface

Formulation/d5796for.pdf

Dental Research.pdf

characteristics and in-vitro behavior of chemically treated bulk Ti6Al7Nb alloys. Surface and Coatings

Technology. 2017;**309**:829-839. DOI: 10.1016/j.surfcoat.2016.10.078

[10] Balanced Hank's Solution. Available from: https://www.sigmaaldrich.com/c ontent/dam/sigma-aldrich/docs/Sigma/

[11] Simulated Saliva Solution. Available from: https://www.pickeringlabs.c om/wp-content/uploads/sds/SDS 1700 0305 Artificial Saliva for Medical and

[12] Dilea M, Mazare A, Ionita D, Demetrescu I. Comparison between corrosion behaviour of implant alloys Ti6Al7Nb and Ti6Al4Zr in artificial saliva. Materials and Corrosion. 2013;

[13] Elias CN, Meirelles L. Improving osseointegration of dental implants. Expert Review of Medical Devices. 2010;**7**:241-256. DOI: 10.1586/erd.09.74

**64**:493-499. DOI: 10.1002/

maco.201206526

[2] Jäger M, Jennissen HP, Dittrich F, Fischer A, Köhling HL. Antimicrobial and osseointegration properties of nanostructured titanium orthopaedic implans. Materials. 2017;**10**:1-28. DOI:

[3] Tamilselvi S, Raman V, Rajendran N. Corrosion behaviour of Ti-6Al-7Nb and Ti-6Al-4V ELI alloys in the simulated body fluid solution by electrochemical

Electrochimica Acta. 2006;**52**:839-846. DOI: 10.1016/j.electacta.2006.06.018

Production of hydroxyapatite layers on the plasma electrolytically oxidized surface of titanium alloys. Materials Science and Engineering: C. 2014;**43**:

[5] Lugovskoy A, Zinigrad M. Plasma electrolytic oxidation of valve metals. In: Mastai Y, editor. Advances in Materials Science and Engineering. London: InTech; 2013. pp. 85-102. DOI: 10.5772/

[6] Ibrahim MZ, Sarhan AAD, Yusuf F, Hamdi M. Biomedical materials and techniques to improve the tribological, mechanical and biomedical properties of orthopedic implants – A review article. Journal of Alloys and Compounds. 2017;

[7] Mohedano M, Matykina E, Arrabal R, Pardo A, Merino MC. Metal release from ceramic coatings for dental implants. Dental Materials. 2014;**30**:e28-e40. DOI:

**714**:636-667. DOI: 10.1016/j.

10.1016/j.dental.2013.12.011

jallcom.2017.04.231

2010;**2**:40-54. DOI: 10.2174/ 1877610801002010040

10.3390/ma10111302

impedance spectroscopy.

527-532. DOI: 10.1016/j. msec.2014.07.030

54827

**154**

[4] Lugovskoy A, Lugovskoy S.

## Biotribology of Mechanically and Laser Marked Biomaterial

*Marcelo de Matos Macedo, Vikas Verma, Jorge Humberto Luna-Domínguez and Ronaldo Câmara Cozza*

#### **Abstract**

The purpose of present work is to study the biotribological behavior of a mechanically and laser marked biomaterial. Sliding wear tests were conducted on ASTM F139 austenitic stainless-steel specimen, with polypropylene and AISI 316 L austenitic stainless-steel balls, as counterbodies. During wear experiments, a liquid chemical composition was continuously fed between the specimen and the ball. The coefficient of friction acting on the tribological system "specimen – liquid chemical composition – ball" and the wear volume of the wear craters were calculated, and results were analyzed. The results have shown that the biotribological behavior of ASTM F139 austenitic stainless steel was influenced by mechanical or laser marking process, and its wear resistance was dependent on the kind of counterbody.

**Keywords:** biomaterial, austenitic stainless steel, laser treatment, wear resistance, wear volume, coefficient of friction

#### **1. Introduction**

The "ball-cratering wear test" has gained large acceptance at universities and research centers as it is widely used in studies focusing on the wear behavior of different materials [1–20]. **Figure 1** presents a schematic diagram of the principle of wear test, where a rotating ball is forced against the tested specimen and liquid solution supplied between the specimen and the ball during the experiments.

The aim of the "ball-cratering wear test" is to generate "wear craters" on the surface of the specimen. **Figure 2** presents an image of such crater, together with an indication of the crater diameter (*d*) (**Figure 2a**) and the wear volume (*V*) (**Figure 2b** [21]). The wear volume is determined as a function of "*d*," using Eq. (1) [22], where *R* is the radius of the ball. *<sup>V</sup>* <sup>≅</sup> *<sup>π</sup> <sup>d</sup>* \_

$$V \cong \frac{\pi d^4}{64R} \text{ for } d \ll \sim R \tag{1}$$

In other line of research, the concept of "biotribology" has gained important spotlight in the area, including research works addressing the biotribological behavior of materials [23–29] used in the manufacturing of human body elements. Consequently, different laboratory techniques have been employed to reproduce conditions where there are friction and consequent wear of parts of the human mechanical structure with relative movement.

However, wear tests conducted under the "ball-cratering" technique present advantages in relation to other types of tribological procedures, as it favors the desired analysis of the tribological behavior. Therefore, considering the need of tribological characterization of biomaterials and the capacity that the "ball-cratering wear test" method presents to this goal, the purpose of this work is to study the biotribological behavior of mechanically and laser-marked ASTM F139 austenitic stainless-steel biomaterial.

**Figure 1.** *"Ball-cratering wear test": representative figure of its operating principle.*

**157**

microscopy.

*Biotribology of Mechanically and Laser Marked Biomaterial*

sition system, in real time, during the sliding wear tests.

Equipment with free-ball mechanical configuration (**Figure 3a**) was used for the sliding wear tests. Two load cells were used in the tribometer: one load cell to control the "normal force – *N*" applied on the specimen (**Figure 3b**) and the other load cell to measure the "tangential force – *T*" developed during the experiments

"Normal" and "tangential" force load cells have a maximum capacity of 50 N and an accuracy of 0.001 N. The values of *N* and *T* were registered by a data acqui-

The tested specimen was an ASTM F139 austenitic stainless-steel biomaterial, marked mechanically and with a nanosecond Q-switched Nd: YAG laser. Its chemi-

Balls of polypropylene and AISI 316 L austenitic stainless steel, with diameter of

To simulate the fluid present in the human body, a chemical liquid solution of PBS – Phosphate Buffered Solution – was inserted between the specimen and the

**Table 3** shows the hardness (*H*) of the materials used in this work (specimen

**Table 4** presents the test conditions defined for the sliding wear experiments

and *D* = 25.4 mm, the tangential sliding velocity (*v*) of the ball is 0.1 m/s. Wear tests were conducted under a test time (*t*) of 10 min. With 0.1 m/s tangential sliding velocity and 10 min (600 s) test time, a sliding distance (*S*) of 60 m was calculated

All experiments were conducted without interruption, and the chemical liquid solution of PBS – Phosphate Buffered Solution – was fed between the specimen and the ball during the tests, under a frequency of 1 drop/10 s. Both the normal force (*N*) and the tangential force (*T*) were monitored and registered constantly. Finalizing the sliding wear tests, the diameters (*d*) of the wear craters were measured by optical microscopy, and their surfaces were analyzed by scanning electron

Finally, the wear volume (*V*) was calculated by Eq. (1), and the coefficient of

*T*

. The rotational speed (*n*) of ball was 75 rpm. For *n* = 75 rpm

and AISI 316 L austenitic stainless

*<sup>N</sup>* (2)

Following values of normal force (*N*) for the sliding wear experiments: *N*PP = 0.05 N and *N*316L = 0.40 N were defined as a function of density (*ρ*) of the

*D* = 25.4 mm (*D* = 1″ – standard size), were adopted as counterbodies.

ball. It was composed by the materials mentioned in **Table 2**.

**2.3 Ball-cratering wear tests and data acquisition**

ball material – polypropylene ⇒*ρ* PP = 0.91 g/cm3

*DOI: http://dx.doi.org/10.5772/intechopen.92564*

**2.1 Ball-cratering wear test equipment**

cal composition is presented in **Table 1**.

**2. Experimental details**

(**Figure 3c**).

**2.2 Materials**

and test balls).

conducted in this research.

between the ball and the specimen.

friction (*μ*) was determined using Eq. (2):

*<sup>μ</sup>* = \_

steel ⇒*ρ* 316L = 8 g/cm3

### **2. Experimental details**

*Corrosion*

**Figure 1.**

stainless-steel biomaterial.

However, wear tests conducted under the "ball-cratering" technique present advantages in relation to other types of tribological procedures, as it favors the desired analysis of the tribological behavior. Therefore, considering the need of tribological characterization of biomaterials and the capacity that the "ball-cratering wear test" method presents to this goal, the purpose of this work is to study the biotribological behavior of mechanically and laser-marked ASTM F139 austenitic

**156**

**Figure 2.**

*Images of wear craters: (a) diameter – d and (b) wear volume – V [21].*

*"Ball-cratering wear test": representative figure of its operating principle.*

#### **2.1 Ball-cratering wear test equipment**

Equipment with free-ball mechanical configuration (**Figure 3a**) was used for the sliding wear tests. Two load cells were used in the tribometer: one load cell to control the "normal force – *N*" applied on the specimen (**Figure 3b**) and the other load cell to measure the "tangential force – *T*" developed during the experiments (**Figure 3c**).

"Normal" and "tangential" force load cells have a maximum capacity of 50 N and an accuracy of 0.001 N. The values of *N* and *T* were registered by a data acquisition system, in real time, during the sliding wear tests.

#### **2.2 Materials**

The tested specimen was an ASTM F139 austenitic stainless-steel biomaterial, marked mechanically and with a nanosecond Q-switched Nd: YAG laser. Its chemical composition is presented in **Table 1**.

Balls of polypropylene and AISI 316 L austenitic stainless steel, with diameter of *D* = 25.4 mm (*D* = 1″ – standard size), were adopted as counterbodies.

To simulate the fluid present in the human body, a chemical liquid solution of PBS – Phosphate Buffered Solution – was inserted between the specimen and the ball. It was composed by the materials mentioned in **Table 2**.

**Table 3** shows the hardness (*H*) of the materials used in this work (specimen and test balls).

#### **2.3 Ball-cratering wear tests and data acquisition**

**Table 4** presents the test conditions defined for the sliding wear experiments conducted in this research.

Following values of normal force (*N*) for the sliding wear experiments: *N*PP = 0.05 N and *N*316L = 0.40 N were defined as a function of density (*ρ*) of the ball material – polypropylene ⇒*ρ* PP = 0.91 g/cm3 and AISI 316 L austenitic stainless steel ⇒*ρ* 316L = 8 g/cm3 . The rotational speed (*n*) of ball was 75 rpm. For *n* = 75 rpm and *D* = 25.4 mm, the tangential sliding velocity (*v*) of the ball is 0.1 m/s. Wear tests were conducted under a test time (*t*) of 10 min. With 0.1 m/s tangential sliding velocity and 10 min (600 s) test time, a sliding distance (*S*) of 60 m was calculated between the ball and the specimen.

All experiments were conducted without interruption, and the chemical liquid solution of PBS – Phosphate Buffered Solution – was fed between the specimen and the ball during the tests, under a frequency of 1 drop/10 s. Both the normal force (*N*) and the tangential force (*T*) were monitored and registered constantly. Finalizing the sliding wear tests, the diameters (*d*) of the wear craters were measured by optical microscopy, and their surfaces were analyzed by scanning electron microscopy.

Finally, the wear volume (*V*) was calculated by Eq. (1), and the coefficient of friction (*μ*) was determined using Eq. (2):

$$
\mu = \frac{T}{N} \tag{2}
$$

#### **Figure 3.**

*(a) "Ball-cratering" wear test equipment with "free-ball" mechanical configuration used for the sliding wear tests: (b) load cell mounted to control the normal force and (c) load cell positioned to measure the tangential force during the experiments.*

**159**

**Table 1.**

**Table 2.**

**Table 3.**

**Table 4.**

**3. Results and discussion**

**3.1 Scanning electron microscopy**

*Hardness of the materials used in this work.*

generated during the sliding wear tests.

**Figure 4** shows a scanning electron micrograph of the surface of a wear crater

*Biotribology of Mechanically and Laser Marked Biomaterial*

**Chemical element % (in weight)** C 0.023 Si 0.78 Mn 2.09 P 0.026 S 0.0003 Cr 18.32 Mo 2.59 Ni 14.33 Fe Balance

*Chemical composition of ASTM F139 austenitic stainless-steel biomaterial – in percentage weight.*

Specimen ASTM F139 austenitic stainless steel 180 HV Test ball Polypropylene 55 – Shore D

Normal force – *N*PP Ball of polypropylene 0.05 N Normal force – *N*316L Ball of AISI 316 L austenitic stainless steel 0.40 N Test ball rotational speed – *n* 75 rpm Tangential sliding velocity – *v* 0.1 m/s Test time – *t* 10 min Sliding distance – *S* 60 m

**Material Hardness –** *H*

AISI 316 L austenitic stainless steel 318 HV

**Chemical element (g/l)** NaCl 8 KCl 0.2 Na2HPO4 1.15 KH2PO4 0.2

*Chemical composition of the PBS – phosphate buffered solution – in g/l.*

*Test parameters for the ball-cratering wear tests under conditions of sliding wear.*

*DOI: http://dx.doi.org/10.5772/intechopen.92564*

#### *Biotribology of Mechanically and Laser Marked Biomaterial DOI: http://dx.doi.org/10.5772/intechopen.92564*


#### **Table 1.**

*Corrosion*

**158**

**Figure 3.**

*force during the experiments.*

*(a) "Ball-cratering" wear test equipment with "free-ball" mechanical configuration used for the sliding wear tests: (b) load cell mounted to control the normal force and (c) load cell positioned to measure the tangential* 

*Chemical composition of ASTM F139 austenitic stainless-steel biomaterial – in percentage weight.*


#### **Table 2.**

*Chemical composition of the PBS – phosphate buffered solution – in g/l.*


#### **Table 3.**

*Hardness of the materials used in this work.*


#### **Table 4.**

*Test parameters for the ball-cratering wear tests under conditions of sliding wear.*

#### **3. Results and discussion**

#### **3.1 Scanning electron microscopy**

**Figure 4** shows a scanning electron micrograph of the surface of a wear crater generated during the sliding wear tests.

Occurrence of grooves, due to sliding movement between the ball and the specimen, was observed in the scanning electron micrograph. The result presented in **Figure 4** is in qualitative agreement with the literature [30], where it is reported that the action of grooves on the surface of a material is characterized as a common tribological behavior of two metallic materials under relative movement.

#### **3.2 Wear volume behavior**

**Figure 5** presents the behavior of the specimen in terms of wear volume (*V*) for the following conditions: mechanically and laser-marked specimen and different types of balls (counterbodies).

#### **Figure 4.**

*Scanning electron micrograph of the surface of a wear crater generated during the sliding wear tests.*

#### **Figure 5.**

*Wear volume (*V*) behavior as a function of the type of marking process ("mechanical" or "laser") and type of counterbody (ball of polypropylene or ball of AISI 316 L austenitic stainless steel). Maximum standard deviation reported: ±7 × 10<sup>−</sup><sup>4</sup> mm3 .*

**161**

**Figure 5**.

**Figure 6.**

*deviation reported: ±0.03.*

**3.3 Coefficient of friction behavior**

different types of balls – counterbodies.

stainless-steel ball counterbodies.

**4. Conclusions**

*Biotribology of Mechanically and Laser Marked Biomaterial*

In addition, **Figure 5** shows a decrease in wear volume under laser marking process for both types of counterbodies. Decrease in wear volume is related to increase in local hardness of the specimen. Increase of hardness can be attributed to the action of laser on specimen surface. In relation to specimen marked mechanically, the possible increase of local hardness could have occurred due to local surface hardening. However, the increase of local hardness caused by laser marking is higher than the local hardness caused by mechanical marking, justifying the results presented in

*Coefficient of friction behavior (*μ*) as a function of the type of marking process ("mechanical" or "laser") and type of counterbody (ball of polypropylene or ball of AISI 316 L austenitic stainless steel). Maximum standard* 

**Figure 6** shows the behavior of the coefficient of friction (*μ*) for the conditions, which the specimen is marked mechanically and marked with laser, and for the

In the present tribological conditions, coefficient of friction was found lower for the wear tests conducted against polypropylene ball than AISI 316 L austenitic

The following conclusions can be drawn from the results obtained in this research, regarding to tribological behavior of ASTM F139 austenitic stainless steel:

• tribological behavior was influenced by the type of the marking process – "mechanical" or "laser" – applied for the investigated biomaterial;

• wear volume was found to be dependent on the normal force acting on the specimen, that is, they were dependent on the type of counterbody – ball of

polypropylene or ball of AISI 316 L austenitic stainless steel; and

*DOI: http://dx.doi.org/10.5772/intechopen.92564*

*Biotribology of Mechanically and Laser Marked Biomaterial DOI: http://dx.doi.org/10.5772/intechopen.92564*

#### **Figure 6.**

*Corrosion*

**3.2 Wear volume behavior**

types of balls (counterbodies).

Occurrence of grooves, due to sliding movement between the ball and the specimen, was observed in the scanning electron micrograph. The result presented in **Figure 4** is in qualitative agreement with the literature [30], where it is reported that the action of grooves on the surface of a material is characterized as a common

**Figure 5** presents the behavior of the specimen in terms of wear volume (*V*) for the following conditions: mechanically and laser-marked specimen and different

tribological behavior of two metallic materials under relative movement.

*Scanning electron micrograph of the surface of a wear crater generated during the sliding wear tests.*

*Wear volume (*V*) behavior as a function of the type of marking process ("mechanical" or "laser") and type of counterbody (ball of polypropylene or ball of AISI 316 L austenitic stainless steel). Maximum standard* 

**160**

**Figure 5.**

*deviation reported: ±7 × 10<sup>−</sup><sup>4</sup>*

 *mm3 .*

**Figure 4.**

*Coefficient of friction behavior (*μ*) as a function of the type of marking process ("mechanical" or "laser") and type of counterbody (ball of polypropylene or ball of AISI 316 L austenitic stainless steel). Maximum standard deviation reported: ±0.03.*

In addition, **Figure 5** shows a decrease in wear volume under laser marking process for both types of counterbodies. Decrease in wear volume is related to increase in local hardness of the specimen. Increase of hardness can be attributed to the action of laser on specimen surface. In relation to specimen marked mechanically, the possible increase of local hardness could have occurred due to local surface hardening.

However, the increase of local hardness caused by laser marking is higher than the local hardness caused by mechanical marking, justifying the results presented in **Figure 5**.

#### **3.3 Coefficient of friction behavior**

**Figure 6** shows the behavior of the coefficient of friction (*μ*) for the conditions, which the specimen is marked mechanically and marked with laser, and for the different types of balls – counterbodies.

In the present tribological conditions, coefficient of friction was found lower for the wear tests conducted against polypropylene ball than AISI 316 L austenitic stainless-steel ball counterbodies.

#### **4. Conclusions**

The following conclusions can be drawn from the results obtained in this research, regarding to tribological behavior of ASTM F139 austenitic stainless steel:


#### *Corrosion*

• coefficient of friction was found dependent on the type of ball; the lower values of *μ* were observed under the use of polypropylene ball.

## **Nomenclature**


**163**

Mexico

**Author details**

Marcelo de Matos Macedo1

and Ronaldo Câmara Cozza1,4\*

Technology (NUST) MISiS, Moscow, Russia

, Vikas Verma2

1 Department of Mechanical Manufacturing, Technology Faculty – FATEC-Mauá, CEETEPS – State Center of Technological Education "Paula Souza", Mauá, SP, Brazil

2 Thermochemistry of Materials SRC, National University of Science and

3 Facultad de Odontología, Universidad Autónoma de Tamaulipas, Tamaulipas,

4 Department of Mechanical Engineering, University Center FEI – Educational

© 2020 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium,

Foundation of Ignatius "Padre Sabóia de Medeiros", SP, Brazil

\*Address all correspondence to: rcamara@fei.edu.br

provided the original work is properly cited.

, Jorge Humberto Luna-Domínguez<sup>3</sup>

*Biotribology of Mechanically and Laser Marked Biomaterial*

*DOI: http://dx.doi.org/10.5772/intechopen.92564*

#### **Greek letters**


*Biotribology of Mechanically and Laser Marked Biomaterial DOI: http://dx.doi.org/10.5772/intechopen.92564*

#### **Author details**

*Corrosion*

**Nomenclature**

**Greek letters**

*H* hardness (HV)

• coefficient of friction was found dependent on the type of ball; the lower

values of *μ* were observed under the use of polypropylene ball.

*d* diameter of the wear crater (mm) *D* diameter of the test ball (mm)

*n* test ball rotational speed (rpm)

*V* wear volume of the wear crater (mm3

*R* radius of the test ball (mm) *S* sliding distance (m) *t* test time (min)

*μ* coefficient of friction ρ density (g/cm3

*N* normal force (applied on the specimen) (N)

*T* tangential force (developed during the wear tests) (N) *v* tangential sliding velocity of the test ball (m/s)

)

)

**162**

Marcelo de Matos Macedo1 , Vikas Verma2 , Jorge Humberto Luna-Domínguez<sup>3</sup> and Ronaldo Câmara Cozza1,4\*

1 Department of Mechanical Manufacturing, Technology Faculty – FATEC-Mauá, CEETEPS – State Center of Technological Education "Paula Souza", Mauá, SP, Brazil

2 Thermochemistry of Materials SRC, National University of Science and Technology (NUST) MISiS, Moscow, Russia

3 Facultad de Odontología, Universidad Autónoma de Tamaulipas, Tamaulipas, Mexico

4 Department of Mechanical Engineering, University Center FEI – Educational Foundation of Ignatius "Padre Sabóia de Medeiros", SP, Brazil

\*Address all correspondence to: rcamara@fei.edu.br

© 2020 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

### **References**

[1] Cozza RC. A study on friction coefficient and wear coefficient of coated systems submitted to micro-scale abrasion tests. Surface & Coatings Technology. 2013;**215**:224-233. DOI: 10.1016/j.surfcoat.2012.06.088

[2] Cozza RC, Rodrigues LC, Schön CG. Analysis of the micro-abrasive wear behavior of an iron aluminide alloy under ambient and high-temperature conditions. Wear. 2015;**330-331**:250- 260. DOI: 10.1016/j.wear.2015.02.021

[3] Cozza RC, JTSL W, Schön CG. Influence of abrasive wear modes on the coefficient of friction of thin films. Tecnologia em Metalurgia, Materiais e Mineração. 2018;**15**(4):504-509. DOI: 10.4322/2176-1523.20181468

[4] Umemura MT, Jiménez LBV, Pinedo CE, Cozza RC, Tschiptschin AP. Assessment of tribological properties of plasma nitrided 410S ferriticmartensitic stainless steels. Wear. 2019;**426-427**:49-58. DOI: 10.1016/j. wear.2018.12.092

[5] Cozza RC. Thin films: Study of the influence of the micro-abrasive wear modes on the volume of wear and coefficient of friction. In: Friction, Lubrication and Wear. 1st ed. London– UK: IntechOpen; 2019. DOI: 10.5772/ intechopen.86459

[6] Cozza RC. Effect of pressure on abrasive wear mode transitions in micro-abrasive wear tests of WC-Co P20. Tribology International. 2013;**57**:266-271. DOI: 10.1016/j.triboint.2012.06.028

[7] Cozza RC. Effect of sliding distance on abrasive wear modes transition. Journal of Materials Research and Technology. 2015;**4**(2):144-150. DOI: 10.1016/j.jmrt.2014.10.007

[8] Wilcken JTSL, Silva FA, Cozza RC, Schön CG. Influence of abrasive wear

modes on the coefficient of friction of thin films. In: Proceedings of the "ICAP 2014 – 2nd International Conference on Abrasive Processes". Cambridge – UK: The University of Cambridge; September 8-10, 2014

[9] Cozza RC. Study of the steady-state of wear in micro-abrasive wear tests by rotative ball conducted on specimen of WC-Co P20 and M2 tool-steel. Revista Matéria. 2018;**23**(1):e-11986. DOI: 10.1590/s1517-707620170001.0322

[10] da Silva WM, Binder R, de Mello JDB. Abrasive wear of steam-treated sintered iron. Wear. 2005;**258**:166-177. DOI: 10.1016/j. wear.2004.09.042

[11] Trezona RI, Allsopp DN, Hutchings IM. Transitions between two-body and three-body abrasive wear: Influence of test conditions in the microscale abrasive wear test. Wear. 1999;**225-229**:205-214. DOI: 10.1016/ S0043-1648(98)00358-5

[12] Adachi K, Hutchings IM. Wearmode mapping for the micro-scale abrasion test. Wear. 2003;**255**:23-29. DOI: 10.1016/S0043-1648(03)00073-5

[13] Adachi K, Hutchings IM. Sensitivity of wear rates in the micro-scale abrasion test to test conditions and material hardness. Wear. 2005;**258**:318-321. DOI: 10.1016/j.wear.2004.02.016

[14] Cozza RC, de Mello JDB, Tanaka DK, Souza RM. Relationship between test severity and wear mode transition in micro-abrasive wear tests. Wear. 2007;**263**:111-116. DOI: 10.1016/j. wear.2007.01.099

[15] Cozza RC, Tanaka DK, Souza RM. Micro-abrasive wear of DC and pulsed DC titanium nitride thin films with different levels of film residual stresses. Surface and Coatings Technology.

**165**

*Biotribology of Mechanically and Laser Marked Biomaterial*

Biomaterialia. 2012;**8**:3888-3903. DOI:

[24] Talha M, Behera CK, Sinha OP. A review on nickel-free nitrogen containing austenitic stainless steels for biomedical applications. Materials Science and Engineering C. 2013;**33**:3563-3575. DOI: 10.1016/j.

[25] Gurappa I. Development of appropriate thickness ceramic coatings on 316 L stainless steel for biomedical applications. Surface and Coatings Technology. 2002;**161**:70-78. DOI: 10.1016/S0257-8972(02)00380-8

[26] Shih CC, Shih CM, Su YY, LHJ S, Chang MS, Lin SJ. Effect of surface oxide properties on corrosion resistance of 316L stainless steel for biomedical applications. Corrosion Science. 2004;**46**:427-441. DOI: 10.1016/ S0010-938X(03)00148-3

[27] Hosseinalipour SM, Ershad-Langroudi A, Hayati AN, Nabizade-Haghighi AM. Characterization of sol–gel coated 316L stainless steel for biomedical applications. Progress in Organic Coatings. 2010;**67**:371-374. DOI: 10.1016/j.porgcoat.2010.01.002

[28] Dewidar MM, Khalil KA, Lim JK. Processing and mechanical properties of porous 3 16L stainless steel for biomedical applications. Transactions of Nonferrous Metals Society of China. 2007;**17**(3):468-473. DOI: 10.1016/

S1003-6326(07)60117-4

[29] Niinomi M. Recent metallic materials for biomedical applications.

Transactions A. 2002;**33A**:477-486. DOI: 10.1007/s11661-002-0109-2

[30] Cozza RC. Third abrasive wear mode: Is it possible? Journal of Materials Research and Technology. 2014;**3**(2):191-193. DOI: 10.1016/j.

Metallurgical and Materials

jmrt.2014.03.010

10.1016/j.actbio.2012.06.037

msec.2013.06.002

*DOI: http://dx.doi.org/10.5772/intechopen.92564*

2006;**201**:4242-4246. DOI: 10.1016/j.

[16] Bose K, Wood RJK. Optimun tests conditions for attaining uniform rolling abrasion in ball cratering tests on hard coatings. Wear. 2005;**258**:322-332. DOI:

surfcoat.2006.08.044

10.1016/j.wear.2004.09.018

[17] Mergler YJ, Huis in 't Veld AJ. Micro-abrasive wear of semi-crystalline polymers. Tribology and Interface Engineering Series. 2003;**41**:165-173. DOI: 10.1016/S0167-8922(03)80129-3

[18] Batista JCA, Matthews A,

S0257-8972(01)01189-6

Godoy C. Micro-abrasive wear of PVD duplex and single-layered coatings. Surface and Coatings Technology. 2001;**142-144**:1137-1143. DOI: 10.1016/

[19] Batista JCA, Godoy C, Matthews A. Micro-scale abrasive wear testing of duplex and non-duplex (singlelayered) PVD (Ti,Al)N, TiN and Cr-N coatings. Tribology International. 2002;**35**:363-372. DOI: 10.1016/ S0301-679X(02)00017-8

[20] Batista JCA, Joseph MC, Godoy C, Matthews A. Micro-abrasion wear testing of PVD TiN coatings on

untreated and plasma nitrided AISI H13 steel. Wear. 2002;**249**:971-979. DOI: 10.1016/S0043-1648(01)00833-X

[21] da Silva WM. Effect of pressing pressure and iron powder size on the micro-abrasion of steam-oxidized sintered iron [M.Sc. Dissertation]. Uberlândia–MG, Brazil: Federal University of Uberlândia; 2003. p. 98

[22] Rutherford KL, Hutchings IM. Theory and application of a micro-scale abrasive wear test. Journal of Testing and Evaluation. 1997;**25**(2):250-260.

[23] Niinomi M, Nakai M, Hieda J. Development of new metallic alloys for biomedical applications. Acta

DOI: 10.1520/JTE11487J

*Biotribology of Mechanically and Laser Marked Biomaterial DOI: http://dx.doi.org/10.5772/intechopen.92564*

2006;**201**:4242-4246. DOI: 10.1016/j. surfcoat.2006.08.044

[16] Bose K, Wood RJK. Optimun tests conditions for attaining uniform rolling abrasion in ball cratering tests on hard coatings. Wear. 2005;**258**:322-332. DOI: 10.1016/j.wear.2004.09.018

[17] Mergler YJ, Huis in 't Veld AJ. Micro-abrasive wear of semi-crystalline polymers. Tribology and Interface Engineering Series. 2003;**41**:165-173. DOI: 10.1016/S0167-8922(03)80129-3

[18] Batista JCA, Matthews A, Godoy C. Micro-abrasive wear of PVD duplex and single-layered coatings. Surface and Coatings Technology. 2001;**142-144**:1137-1143. DOI: 10.1016/ S0257-8972(01)01189-6

[19] Batista JCA, Godoy C, Matthews A. Micro-scale abrasive wear testing of duplex and non-duplex (singlelayered) PVD (Ti,Al)N, TiN and Cr-N coatings. Tribology International. 2002;**35**:363-372. DOI: 10.1016/ S0301-679X(02)00017-8

[20] Batista JCA, Joseph MC, Godoy C, Matthews A. Micro-abrasion wear testing of PVD TiN coatings on untreated and plasma nitrided AISI H13 steel. Wear. 2002;**249**:971-979. DOI: 10.1016/S0043-1648(01)00833-X

[21] da Silva WM. Effect of pressing pressure and iron powder size on the micro-abrasion of steam-oxidized sintered iron [M.Sc. Dissertation]. Uberlândia–MG, Brazil: Federal University of Uberlândia; 2003. p. 98

[22] Rutherford KL, Hutchings IM. Theory and application of a micro-scale abrasive wear test. Journal of Testing and Evaluation. 1997;**25**(2):250-260. DOI: 10.1520/JTE11487J

[23] Niinomi M, Nakai M, Hieda J. Development of new metallic alloys for biomedical applications. Acta

Biomaterialia. 2012;**8**:3888-3903. DOI: 10.1016/j.actbio.2012.06.037

[24] Talha M, Behera CK, Sinha OP. A review on nickel-free nitrogen containing austenitic stainless steels for biomedical applications. Materials Science and Engineering C. 2013;**33**:3563-3575. DOI: 10.1016/j. msec.2013.06.002

[25] Gurappa I. Development of appropriate thickness ceramic coatings on 316 L stainless steel for biomedical applications. Surface and Coatings Technology. 2002;**161**:70-78. DOI: 10.1016/S0257-8972(02)00380-8

[26] Shih CC, Shih CM, Su YY, LHJ S, Chang MS, Lin SJ. Effect of surface oxide properties on corrosion resistance of 316L stainless steel for biomedical applications. Corrosion Science. 2004;**46**:427-441. DOI: 10.1016/ S0010-938X(03)00148-3

[27] Hosseinalipour SM, Ershad-Langroudi A, Hayati AN, Nabizade-Haghighi AM. Characterization of sol–gel coated 316L stainless steel for biomedical applications. Progress in Organic Coatings. 2010;**67**:371-374. DOI: 10.1016/j.porgcoat.2010.01.002

[28] Dewidar MM, Khalil KA, Lim JK. Processing and mechanical properties of porous 3 16L stainless steel for biomedical applications. Transactions of Nonferrous Metals Society of China. 2007;**17**(3):468-473. DOI: 10.1016/ S1003-6326(07)60117-4

[29] Niinomi M. Recent metallic materials for biomedical applications. Metallurgical and Materials Transactions A. 2002;**33A**:477-486. DOI: 10.1007/s11661-002-0109-2

[30] Cozza RC. Third abrasive wear mode: Is it possible? Journal of Materials Research and Technology. 2014;**3**(2):191-193. DOI: 10.1016/j. jmrt.2014.03.010

**164**

*Corrosion*

**References**

[1] Cozza RC. A study on friction coefficient and wear coefficient of coated systems submitted to micro-scale abrasion tests. Surface & Coatings Technology. 2013;**215**:224-233. DOI: 10.1016/j.surfcoat.2012.06.088

modes on the coefficient of friction of thin films. In: Proceedings of the "ICAP 2014 – 2nd International Conference on Abrasive Processes". Cambridge – UK: The University of Cambridge;

[9] Cozza RC. Study of the steady-state of wear in micro-abrasive wear tests by rotative ball conducted on specimen of WC-Co P20 and M2 tool-steel. Revista Matéria. 2018;**23**(1):e-11986. DOI: 10.1590/s1517-707620170001.0322

[10] da Silva WM, Binder R, de Mello JDB. Abrasive wear of steam-treated sintered iron. Wear. 2005;**258**:166-177. DOI: 10.1016/j.

[11] Trezona RI, Allsopp DN, Hutchings IM. Transitions between two-body and three-body abrasive wear: Influence of test conditions in the microscale abrasive wear test. Wear. 1999;**225-229**:205-214. DOI: 10.1016/

S0043-1648(98)00358-5

10.1016/j.wear.2004.02.016

wear.2007.01.099

[14] Cozza RC, de Mello JDB,

Tanaka DK, Souza RM. Relationship between test severity and wear mode transition in micro-abrasive wear tests. Wear. 2007;**263**:111-116. DOI: 10.1016/j.

[15] Cozza RC, Tanaka DK, Souza RM. Micro-abrasive wear of DC and pulsed DC titanium nitride thin films with different levels of film residual stresses. Surface and Coatings Technology.

[12] Adachi K, Hutchings IM. Wearmode mapping for the micro-scale abrasion test. Wear. 2003;**255**:23-29. DOI: 10.1016/S0043-1648(03)00073-5

[13] Adachi K, Hutchings IM. Sensitivity of wear rates in the micro-scale abrasion test to test conditions and material hardness. Wear. 2005;**258**:318-321. DOI:

wear.2004.09.042

September 8-10, 2014

[2] Cozza RC, Rodrigues LC, Schön CG. Analysis of the micro-abrasive wear behavior of an iron aluminide alloy under ambient and high-temperature conditions. Wear. 2015;**330-331**:250- 260. DOI: 10.1016/j.wear.2015.02.021

[3] Cozza RC, JTSL W, Schön CG. Influence of abrasive wear modes on the coefficient of friction of thin films. Tecnologia em Metalurgia, Materiais e Mineração. 2018;**15**(4):504-509. DOI:

10.4322/2176-1523.20181468

wear.2018.12.092

intechopen.86459

[4] Umemura MT, Jiménez LBV,

Pinedo CE, Cozza RC, Tschiptschin AP. Assessment of tribological properties of plasma nitrided 410S ferriticmartensitic stainless steels. Wear. 2019;**426-427**:49-58. DOI: 10.1016/j.

[5] Cozza RC. Thin films: Study of the influence of the micro-abrasive wear modes on the volume of wear and coefficient of friction. In: Friction, Lubrication and Wear. 1st ed. London– UK: IntechOpen; 2019. DOI: 10.5772/

[6] Cozza RC. Effect of pressure on abrasive wear mode transitions in

micro-abrasive wear tests of WC-Co P20. Tribology International. 2013;**57**:266-271. DOI: 10.1016/j.triboint.2012.06.028

[7] Cozza RC. Effect of sliding distance on abrasive wear modes transition. Journal of Materials Research and Technology. 2015;**4**(2):144-150. DOI:

[8] Wilcken JTSL, Silva FA, Cozza RC, Schön CG. Influence of abrasive wear

10.1016/j.jmrt.2014.10.007

## *Edited by Ambrish Singh*

Corrosion is a global threat and a burning topic for new and innovative research. Corrosion causes shut downs, economic losses, delays, failures, accidents, losses of life, and losses in productivity. "Wherever metal is, there corrosion will occur" – this is a general concept as not many protection methods are available to mitigate corrosion. The available methods can only delay the process but cannot stop or protect the metal completely. So there is always a need for good research and inventions in this field. This book includes the recent research work done in the field of corrosion. The chapters are written by reputed authors in the field of corrosion and have been reviewed extensively before acceptance. The chapters focus on different aspects of corrosion to provide readers with a good idea of the overall process. The diversification of the chapters will keep the readers interested and motivated for new innovations in the field of corrosion. It will be very useful to scholars, academicians, researchers, and industrialists.

Published in London, UK © 2020 IntechOpen © holwichaikawee / iStock

Corrosion

Corrosion

*Edited by Ambrish Singh*